This invention relates to a system for deriving an intraocular lens (IOL), suited to a particular patient and for replacing their natural but diseased crystalline lens, and more specifically, to a system for deriving the power and other parameters defining the IOL, based on corneal surface measurements (both anterior and posterior) and ocular length measurements (lens thickness, lens position, and axial length) of the patient""s eye.
Everyone, if they live long enough, will develop cataracts, which are degenerative opacities in the normal crystalline lens that restrict vision and hence quality of life. The normal course of treatment is to surgically remove the cataractous lens and replace it with a synthetic intraocular lens (IOL). Although there are many IOL varieties, their function is the same: to bring images seen by a person into sharp focus on their retina without the aid of other forms of correction (spectacles or contact lenses). To do this, the IOL must have the proper optical power. Excessive power causes the image to form in front of the retina, while insufficient power causes the image to form behind the retina. Only when the image focuses on the retina can the image appear sharp.
There are two major classes of IOLs in current use: spheric and toric. Toric IOLs can correct for astigmatism, which is the optical aberration characterized by a 2-fold sinusoidal variation of power with meridional angle. Persons having xe2x80x9cwith the rulexe2x80x9d astigmatism have their greatest power aligned primarily in their infero-superior (i.e., vertical) plane, while persons having xe2x80x9cagainst the rulexe2x80x9d astigmatism have greatest power aligned primarily in their naso-temporal (i.e., horizontal) plane. Fewer patients have oblique alignments lying between the two more common extremes.
Spheric IOLs are characterized by a single power and can not correct for astigmatism. A third class of specialty IOLs is in limited clinical use and combines unusual features (like multi-focal power) to overcome special ocular defects (like presbyopia). A forth class of customized IOLs is not in clinical use but potentially could correct for higher order optical aberrations. There is a significant difference between specialty and customized IOLs. The latter attempts to provide a perfect optical image for a particular ocular state, while the former compromises image quality in an attempt to accommodate differing states of the eye.
Many factors contribute to the image that is delivered to and focused on the retina. Portions of the eye affecting this image include the cornea, the crystalline lens, the aqueous and vitreous humors within the eye, retinal shape, and the superficial tear film covering the cornea. In addition, optical focus is affected by both the shape (curvature, in particular) and location of all internal refracting surfaces, the refractive indices of the intervening material, as well as the axial length of the eye as a whole.
Much work has been done by many researchers to provide both theoretical and empirical formulas for calculating the IOL power suited to a particular patient. Clinically established IOL power formulas include SRK, SRK II, SRK/T, Holladay I, Holladay II, Binkhorst, Olsen, Hoffer-Calenbrander, TMB, DKG, and WPC formulas. Retzlaff, et al. in their book, Lens Implant Power Calculation, extensively discuss many lens power calculation formulas and the factors influencing the power calculation.
The IOL power calculation formulas listed above take into account many factors such as corneal curvature, anterior chamber depth, corneal size in terms of horizontal-white-to-white distance, anterior chamber depth, crystalline lens thickness, and the axial length of the eye. However, these formulas also estimate corneal power solely from measurements of the keratometric curvature (K) of the anterior corneal surface, which is a single number (or two in the case of astigmatism) typically measured by a keratometer. No information concerning the actual posterior corneal power is included in these formulas.
The details of how K is employed vary from formula to formula. For example in SRK, K is just a parameter in a regression analysis of the data. Holladay, on the other hand, tries to estimate the actual corneal power from K. This is done by assuming the posterior radius is exactly 1.2 mm smaller than the anterior radius deduced from K (see the Journal of Cataract and Refractive Surgery, volume 23, pages 1360-1361, November 1997). Although Holladay""s approach may seem more satisfying, in that posterior power is not ignored, no more information is really added. Because empirical information is always used to tune each formula, the effect of posterior power is included automatically, but only for the normal, population-averaged, cornea.
Not accounting for the actual power of posterior cornea introduces the potential for error in the IOL derivation. Only when the posterior cornea closely follows the anterior cornea, which admittedly constitutes the majority of cases, is this neglect tolerable. However, when the posterior cornea diverges significantly from its anterior surface, the final surgical result is a xe2x80x9crefractive surprisexe2x80x9d that could only have been anticipated had the posterior corneal surface been measured and its power properly taken into account.
Another simplification that introduced error in the power calculation is not accounting for localized wave speed, interface refraction, and alignment of the optical and acoustic probe beams employed in ocular measuring devices. In humans, the line of sight is not generally aligned along the optical axis of the eye, but is often skewed several degrees from the optical axis. Light and sound probe beams not only refract at the tilted interfaces, but refract in opposite directions. This occurs because denser materials typically slow light while increasing sound speed. Also A-scan ultrasound instruments are typically aligned along the optical axis (giving the largest reflections), while optical instruments are aligned along the visual axis.
Another error-inducing simplification occurs when the retina is approximated as flat or uniformly receptive. Although a flat image plane is consistent with paraxial optics calculations, more sophisticated analyses that simulate the extended image (or point spread function) are affected by the curvature inherent in the photoreceptor surface. Spherical aberration is particularly affected. The long narrow shape of the photoreceptors gives them an angle dependent receptivity, not unlike a fiber optic. When taken as a whole, this is quantified by the well-known Stiles-Crawford effect.
A major failing of current technology is its inability to calculate IOL powers for eyes with non-normal corneas. These include all eyes compromised by corneal disease and corneal and refractive surgery. The fundamental problem stems from trying to assign a single curvature value (or two in the case of astigmatism) to a non-spherical surface. Keratometric curvature, the second most important parameter in current IOL power formulas, makes this assignment assuming the anterior corneal surface is spherical in shape. In actuality, the normal virgin cornea is prolate ellipsoidal, a difference whose consequences are hidden within the empirical constructs of these formulas. Post-operative corneas have distinctly different shapes, and therefore, they confound previous empiricisms. Such misshapen surface errors can only be corrected if the actual corneal shape (not just its keratometric curvature) is measured and properly taken into account.
Therefore, there is a need for a system where the actual measured corneal surfaces, both anterior and posterior, and the alignments, shapes, and positions of these and other important ocular components are used collectively in the derivation of the IOL that best fits a patient""s eye.